It is common practice today to use mechanical devices to maintain a supply of oxygen and blood circulation during heart failure. One type of device used for this purpose is known as the heart-lung machine. A heart-lung machine performs the function of the heart and lungs. Venous blood is removed from the patient's body and passed through an external oxygenator. The oxygenated blood is then returned to the patient's body. Heart-lung machines are frequently used during open-heart surgery to maintain a supply of oxygen and blood circulation.
Another method which is used to maintain blood circulation is a ventricular actuation cup. These devices typically comprise a cup which fits around the human heart. A thin membrane is bonded to the inner surface of the cup and functions as a diaphragm. Positive and negative pressure is applied to the space between the membrane and the cup wall to alternately inflate and deflate the diaphragm. When the diaphragm is inflated, the heart is squeezed to simulate systolic action. Similarly, when the diaphragm is deflated, the heart muscle relaxes to simulate diastolic action.
The ventricular actuation cups described above have been proven effective in the treatment of patients subject to heart failure. The blood circulation of such patients can be maintained for periods up to several weeks with such devices. However, ventricular actuation cups have some drawbacks.
One frequently encountered problem with the present ventricular actuation cups currently in use relates to the possibility of the failure of the diaphragm. In prior art designs, the diaphragm is bonded to the inner surface of the cup. During the systolic pumping phase, the diaphragm is progressively distended. Several stress factors localized at the bond zones are significantly increased in magnitude. The stress factor associated with a lifting or tearing force at the bond zone reaches its maximum value at maximum systolic distention. This lifting/tearing vector tends to separate the membrane from the cup wall. Continued cyclic development of this stress vector may lead to bond failure which renders the device inoperative. When a bonding failure occurs, the patient's life becomes seriously threatened.
Another significant drawback of currently available ventricular actuation devices is that right and left ventricle pumping is accomplished by using a single diaphragm. As a result, it is difficult to simulate the normal heart function. It is well established that differences exist between right and left ventricle output. It is known, for instance, that right and left ventricular pressures are different. Normal ventricle pressures are 120 over 12-15 mm Hg for the left ventricle and 25-30 over 0-5 mm Hg for the right ventricle. As a result, the pressure selected for ventricular actuation will, at best, be a compromise between the optimum pressure for either the left or right ventricle. Any adjustment of-left ventricular flow or pressure will alter right ventricular pumping parameters. Thus, it is virtually impossible to achieve optimum output and pressures for both the right and left ventricle using a single diaphragm.
The design of currently available ventricular actuation cups also leads to trauma of tissue and organs surrounding the heart. Most cups have a large port molded to the outer surface of the cup which is used for actuation of the diaphragm. When the cup is positioned to contain the heart, the combined reaction to the continuous systolic/diastolic pumping action causes the cup to continually oscillate about its longitudinal axis. The magnitude of this longitudinal oscillation can vary from as much as +15.degree. to -15.degree.. As the device oscillates, the relatively high profile inlet port rubs against surrounding tissue and lungs. The continuous rubbing or abrasion of lung tissue may lead to progressive development of atelectasis in which the alveolar cells in the traumatized lung tissue collapse. In the most serious cases, the continuous trauma has resulted in pulmonary bleeding with resultant pulmonary function compromise.
Another consideration that existing ventricular actuation cups often fail to take into account is the variation in size of the human heart from one person to another. Furthermore, it is possible that different size cups may be needed for the same patient. Certain physiological conditions, such as pulmonary hypertension can result in dilation of the heart. When the ventricular actuation cup is initially placed, a relatively large cup may be needed to accommodate the dilated heart. As the heart returns to a more normal size, the efficiency of the oversize cup in maintaining blood circulation may diminish to the point that a smaller cup must be used. This problem is usually addressed by manufacturing cups of different size to accommodate different size hearts. The surgeon must therefore have a relatively large number of differently sized cups on hand to select from depending on the size of his patient's heart and the patient's physiological condition.
Current univentricular devices, e.g. Anstadt, U.S. Pat. No. 5,119,804 (incorporated herein by reference), presently use a rotary vane combination pressure-vacuum pump as the common source for obtaining and maintaining the pressures and vacuums required for systolic and diastolic ventricular cup actuation. Systolic and diastolic pumping action times are essentially controlled by a combination of timers operating solenoid valves in both the pressure and vacuum lines such that the single drive line delivers a timed sequence of diaphragm distending systolic pressure pulses and diaphragm retraction diastolic vacuum pulses.
An unavoidable requirement in this type of drive line source is a mechanical device called a "damper valve" in the primary drive line exiting the sequencing pressure-vacuum manifold. This valve serves the primary function of "damping" the transition from systolic-diastolic pumping action, in order to prevent diaphragm "slap" during diastolic relaxation. The damper valve is also used in establishing the pressure profile of systolic pumping action e.g. applying during early systolic compression approximately 25-30 mm.Hg and during late systolic compression approximately 120-150 mm Hg. Similarly, during the diastolic phase of each cycle, vacuum pressures of approximately -100 to -120 mm Hg are applied.
The requirement of the damping valve for purposes just set forth totally prohibits simultaneous and independent control of both left and right ventricular pumping action, in terms of required physiological flow rates and pressures, and related to patient forward flow requirements of a changing nature with time.